1. Field Of The Invention
This invention relates generally to the field of magnetic resonance imaging systems and, more specifically, to the transmitter in a magnetic resonance imaging system which transmitter contains a pulse control system to aid in the production of a radio-frequency current pulse in the transmitter coil with a desired magnitude to improve imaging results.
2. Description Of Related Art
Magnetic resonance imaging ("MRI"), also known as nuclear magnetic resonance ("NMR") imaging, has become a valuable tool as a safe, non-invasive means for obtaining information in the form of images of objects under examination. For example, MRI can be used as a medical diagnostic tool by providing images of selected portions of the human body without the use of X-ray photography.
In such an MRI system, a transmitter system is utilized to generate a high-frequency magnetic field within the imaging volume of the object under examination in order to manipulate the spins of the relevant nuclei within the imaging volume as is well known in the art. A typical MRI transmitter system shown in a basic configuration is illustrated in FIG. 1. MRI transmitter 10 includes radio-frequency ("RF") pulse generator 11, transmitter amplifier 12, coupling network 13, and transmitter coil 14. RF pulse generator 11 produces a modulated carrier output on its output 15 of desired shape and magnitude. Typically, the RF pulse produced is a voltage waveform having a repeated rectangular pulse shape or a sinc pulse shape of desired time duration and magnitude. The RF pulse is produced by a modulator which has as inputs an RF carrier and a desired modulating input. The RF pulses are fed into input 16 of linear transmitter amplifier 12 and amplified from a typical 0.1-1.0 watt signal to pulses having a power in the range of 100 to 10,000 watts on the amplifier's output 17. The amplifier should provide a linear response over the operating frequency range of the transmitter.
Coupling network 13 receives the amplified RF pulse from amplifier 12 and produces the desired RF current in transmitter coil 14. That current, in turn, generates the desired high-frequency magnetic field inside the transmitter coil. Coupling network 13 usually consists of a tuned transformer network which acts to resonate transmitter coil 14 and to provide a proper load impedance for the output of transmitter amplifier 12. Typical operating frequencies range from less than 1 megahertz to 100 megahertz and more typically range between 5 and 80 megahertz. The desired RF current is dependent upon the specific MRI system, the material being examined, the type of information sought from the scanning and the like and is readily determined by a person of ordinary skill in the art.
For proper operation of the imager of the MRI system, the generated high-frequency magnetic field must be a pulse with a definite time-dependence (that is, shape) and definite absolute value (that is, magnitude) to manipulate properly the spins of the relevant nuclei within the imaging volume. For example, a pulse of specified duration and field strength, sinusoidally oscillating at the Larmor resonance frequency for the nuclear spin, can be applied to the volume in order to rotate the net nuclear spin magnetization of the relevant nuclei of the patient or object under examination in a desired manner for proper imaging.
In practice, the time duration of the high-frequency magnetic field pulse can be easily controlled. By using a transmitter amplifier, such as amplifier 12 in FIG. 1, the shape of the pulse can also be controlled. The magnitude of the actual high-frequency magnetic field within the imaging volume will depend mainly on the RF current through the transmitter coil, such as coil 14 of transmitter 10 in FIG. 1. The relationship between this RF current and the RF pulse input to the transmitter amplifier is sensitive, however, to the details of the coupling between transmitter coil 14 and the patient or object under examination since the latter will affect both the reactance ("detuning") and resistance ("loading") of the transmitter coil at the operating frequency. The current through coil 14 produced from the RF pulses out of transmitter amplifier 12 is, therefore, a function of the coil impedance which is affected by the above discussed reactance and resistance. Thus, the current will vary with a change in position, size, conductivity, etc. of the patient or object under examination or motion of the patient within the imaging volume.
Ideally, more accurate control of the high-frequency magnetic field would be achieved by detection of the actual magnetic field value in the imaging volume during scanning of an object under examination and by adjusting the RF pulses to produce the desired magnetic field value. However, that process is not practical since the most appropriate location for monitoring that field is already occupied by the object under examination during scanning. A more practical method of attempting to obtain a consistent and appropriate magnitude of the high-frequency magnetic field is to produce an RF pulse in the MRI transmitter with an object in the imaging volume, receive the resulting NMR signal in the MRI system's receiver, and then maximize the received signal by adjusting the gain of the transmitter amplifier in the MRI transmitter. This method, however, is a time consuming one for optimizing the RF pulses and cannot be done while imaging, that is, during scanning. Furthermore, as discussed above, coupling may change the current in the transmitter coil due to the change of the object under examination or due to movement of that object during examination.
To obtain a consistent, desired magnitude of the high-frequency magnetic field from one imaging scan to another, it would be beneficial to adjust the transmitter amplifier gain to obtain a consistent, desired magnitude of the coil current from scan to scan. This would be much more convenient than adjusting the actual magnetic field value and would be much faster then the method described above. In principle, the RF current through transmitter coil 14 could be stabilized by applying negative feedback to transmitter amplifier 12 from a sample of the current through the transmitter coil. In practice, however, the sine-wave carrier is a high-frequency carrier (typically 5 to 80 megahertz), and transmitter coil 14 is located far from transmitter amplifier 12 (usually more than 10 meters away). Thus, there would be, in practice, a phase shift caused by the propagation delay through the transmitter output and feedback sample cables. That unwanted phase shift makes direct feedback at the carrier frequency extremely difficult and impractical. For example, the use of 20 meters of coaxial cable having a velocity factor of 0.66 will give 180.degree. of phase shift at approximately 5 megahertz operating frequency. Furthermore, attempts at stabilizing the RF current out of the transmitter amplifier are not sufficient since the ratio between the amplifier's output current and the coil current is affected by the detuning and loading caused by the object being examined as discussed above.
From the foregoing considerations, it should be apparent that there is a need for an improved MRI transmitter system in which an RF current pulse in the transmitter coil can be controlled to maintain a desired magnitude during scanning of an object under examination.
It is, thus, intended that the invention provide a pulse control system for an MRI transmitter in which there is improved performance.
Another intent is that the invention provide a pulse control system to produce an RF current pulse in the transmitter coil with a desired magnitude during scanning.
Still another intent is that the invention provide a pulse control system for an MRI transmitter to compensate for tuning and loading changes in the transmitter coil caused by variations in size, position, and conductivity of the object under examination.
Yet another intent is that the invention provide a pulse control system for an MRI transmitter to correct the magnitude of the RF current in the transmitter coil to a desired magnitude which correction avoids the problems of phase shift at carrier frequency.
Other intentions and features of the invention will further become apparent with reference to the accompanying drawings and the detailed description of the invention or may be learned by practice of the invention.